Medical imaging is one of the most useful diagnostic tools available in modern medicine. Medical imaging allows medical personnel to non-intrusively look into a living body in order to detect and assess many types of injuries, diseases, conditions, etc. Medical imaging allows doctors and technicians to more easily and correctly make a diagnosis, decide on a treatment, prescribe medication, perform surgery or other treatments, etc.
There are medical imaging processes of many types and for many different purposes, situations, or uses. They commonly share the ability to create an image of a bodily region of a patient, and can do so non-invasively. Examples of some common medical imaging types are nuclear medical (NM) imaging such as positron emission tomography (PET) and single photon emission computed tomography (SPECT), electron-beam X-ray computed tomography (CT), magnetic resonance imaging (MRI), and ultrasound (US). Using these or other imaging types and associated machines, an image or series of images may be captured. Other devices may then be used to process the image in some fashion. Finally, a doctor or technician may read the image in order to provide a diagnosis.
In positron emission tomography (PET), the detector needs to have a high efficiency (energy range of 511 keV), and it needs to detect this radiation with sufficient energy resolution (10-15% FWHM) to distinguish unscattered from scattered and background radiation. The detector also needs to count single events and provide a very accurate timing resolution of few ns or better, in order to identify coincidences between pairs of 511 keV events that originate from the same positron decay. In addition, a spatial resolution of a few mm is required for clinical systems and in the <1 mm range for pre-clinical applications.
State-of-the-art PET detectors are mostly based on scintillators and photosensors. The scintillator converts the high-energy gamma radiation into visible light, then the photodetector converts the visible photons into an electrical signal, which is usually further amplified by the front-end readout electronics.
The standard detector design today is a block detector, in which an array of discrete scintillator crystals is viewed by a smaller number of photomultiplier tubes (PMTs). This optical multiplexing leads to a strong reduction of the number of sensor channels and electronics channels, compared to the number of scintillator crystals. The event position then can be obtained by calculating the centroid of the detected PMT signals and assigning the event to one of the discrete crystals by a lookup map.
Another approach uses a large continuous scintillator crystal with an array of PMTs, where the event positions are determined by classic Anger centroid calculation from the PMT signals, similar to a SPECT detector. Nal(Tl) scintillator plates are mostly used for such designs. The drawback of this approach is that the positioning becomes less accurate at areas of high crystal thickness, especially near the edges of the scintillator plate. In the case of Nal(Tl), the timing resolution is relatively poor due to the long decay time, and the pile-up of simultaneous events on the scintillator plate becomes problematic at higher count rates.
The present invention provides a high-resolution PET detector, which simultaneously provides good energy resolution, good timing and accurate position information up to the scintillator edges.